Aberrations in optical microscopy reduce image resolution and contrast, and can limit imaging depth when focusing into biological samples, for example. Static correction of aberrations may be achieved through appropriate lens design, but this does not offer the flexibility of simultaneously correcting aberrations for all imaging depths, nor the adaptability to correct for sample-specific aberrations for high-quality tomographic optical imaging.
It has proven possible to improved optical image contrast and resolution considerably by applying adaptive-optics (AO) methods, however these methods have been limited to microscopic techniques based either on intensity (non-interferometric) imaging (as taught, for example, by Wright, et al., “Adaptive optics for enhanced signal in CARS microscopy,” Opt. Express, vol. 15, pp. 18209-19 (2007), or on optical coherence tomography (OCT), as taught, for example, by Hermann, et al., “Adaptive-optics ultrahigh-resolution optical coherence tomography,” Opt. Lett., vol. 29, pp. 2142-44 (2004), and Zhang, et al., “High-speed volumetric imaging of cone photoreceptors with adaptive optics spectral-domain optical coherence tomography,” Opt. Express, vol. 14, pp. 4380-94 (2006), all of which papers are incorporated herein by reference. Booth et al., Predictive aberration correction for multilayer optical data storage, Appl. Phys. Lett., vol. 88, 031109 (2006), incorporated herein by reference, suggests the application of beam-correcting AO techniques to three-dimensional optical data storage devices. All of the foregoing methods operate by correcting a beam used in deriving an image.
The light microscope is a fundamental tool underpinning many historic developments in medicine and biology. The confocal laser scanning microscope utilizes a pinhole to reject light from out-of-focus planes to achieve superior optical sectioning, and enable 3D imaging (tomography). Modern optical microscopy, capitalizing on the development of the laser, has provided new capabilities to image thick specimens, and peer deeper into scattering tissues. Additionally, two-photon microscopy advantageously provides for imaging depths of hundreds of micrometers in biological tissue.
The development of OCT has enabled in vivo tomography with a relatively large imaging depth (1-3 mm) in scattering tissues, as discussed, for example, in Fercher et al., Optical coherence tomography—principles and applications, Rep. Prog. Phys., vol. 66, pp. 239-303 (2003), which is incorporated herein by reference. OCT has achieved widespread clinical use in ophthalmology, and applications in cardiology, oncology, gastroenterology, and dermatology are currently undergoing translation from the research lab into clinical practice. By incorporating near-infrared broadband sources, OCT can provide micrometer-scale axial resolution. However, transverse resolution, which is inversely proportional the objective lens numerical aperture (NA), is typically low, resulting in an asymmetric three-dimensional (3D) point-spread function (PSF). The use of higher NA optics has enabled cellular resolution, as discussed in Boppart et al., In vivo cellular optical coherence tomography imaging, Nat. Med., vol. 4, pp. 861-65 (1998), incorporated herein by reference, however the use of high-NA optics results in significant reduction of the depth-of-field.
A solution to the problem of reduced depth of field has been that of combining tomograms obtained at different focal depths, as taught by Rolland et al., Gabor-based fusion technique for Optical Coherence Microscopy, Opt. Exp., vol. 18, 3632-42 (2010), incorporated herein by reference, however the price of acquiring and combining multiple tomograms is increased acquisition time and the complication of mechanical scanning. Another solution is that of Interferometric synthetic aperture microscopy (ISAM), taught in Ralston et al., Interferometric synthetic aperture microscopy, Nat. Phys., vol. 3, pp. 129-34 (2007), incorporated herein by reference, a computed imaging technique based on a solution to the inverse scattering problem for OCT, enables object reconstruction with spatially invariant focal-plane resolution at all depths, without having to scan the focus. The foundations of interferometric synthetic aperture microscopy (ISAM) are laid out in U.S. Pat. No. 7,602,501 (issued Oct. 13, 2009) and U.S. Pat. No. 7,643,155 (issued Jan. 5, 2010), and in references cited in each of the foregoing patents, all of which patents and references are incorporated herein by reference. ISAM provides for the scattering potential of a sample to be reconstructed with spatially invariant resolution even outside a region of focus, using modalities such as optical coherence tomography (OCT). Thus, deep regions may be imaged with high transverse resolution even with optics of high numerical aperture (NA), where the depth of focus is short.
A remaining problem, however, even when shortened depth of focus at high-NA is effectively remediated through the inverse scattering techniques of ISAM, is that of aberrations introduced by the optical system and by characteristics of the sample itself.
Aberration correction has been demonstrated in digital holographic microscopy (DHM), as described, for example, by Colomb et al., Numerical parametric lens for shifting, magnification, and complete aberration compensation in digital holographic microscopy, J. Opt. Soc. Am. A, vol. 23, 3177-90 (2006), and in De Nicola et al., Recovering correct phase information in multiwavelength digital holographic microscopy by compensation for chromatic aberrations, Opt. Lett., vol. 30, 2706-08 (2005), both of which of are incorporated herein by reference. Computational correction of aberrations in DHM, however, has only been demonstrated for non-scattering (or, “thin”) samples using discrete-wavelength optical sources, as by Miccio, et al., Direct full compensation of the aberrations in quantitative phase microscopy of thin objects by a single digital hologram, Appl. Phys. Lett., vol. 90, 041104 (2007). DHM's susceptibility to cross-talk from adjacent regions of a sample and to out-of-plane scattering in turbid samples has rendered it ineffective for tomography of bulk biological tissue.
A method of Kam et al., Computational adaptive optics for live three-dimensional biological imaging, Proc. Natl. Acad. Sci. USA, vol. 98, pp. 3790-95 (2001), uses space-variant deconvolution to compensate sample-specific aberrations, and is suited for imaging fluorescence from relatively weakly scattering biological samples. Demonstrated with an image of a fluorescent bead under an oil droplet, this method utilized a separate measurement of the sample using Nomarski differential interference microscopy to map its refractive index. This map was then used to perform 3D optical ray tracing to compute the magnitude of the aberrated PSF used for deconvolution.
None of the methods taught in the prior art, however, has been applicable to interferometric tomographic techniques wherein spectral information is pertinent to data analysis.